There already exists a device to display disintegrations of positrons; the positron tomograph which uses the characteristic of the positrons which disintegrate when encountering an electron by emitting two gamma radiations directed alond two opposing directions, each radiation having an energy of 511 keV.
FIG. 1 diagrammatically shows a cutaway view of a positron tomograph.
During a preliminary stage, a substance containing radioactive elements emitting positrons is injected into the organ to be displayed 10 (in this instance, a brain with living tissue, but this organ can be mechanical). The organ 10 is placed at the center of a ring 12 of detectors 14, each formed of a scintillator sensitive to gamma radiations and coupled to a photoelectron multiplier. Each positron emitted disintegrates with one electron after an average free distance covered depending on the organ 10 studied (this average free distance covered is between 2 and 3 mm in the human body). A disintegration gives rise to an emission of two gamma radiations 11 in opposing directions which are detected in coincidence by the detectors 14 of the ring. These coincidence detections embody an electronic collimation which makes it possible to materialize the track of the coupled gamma radiations.
Next, by means of an image reconstruction treatment using operations such as retroprojections of tracks along several directions, convolutions, etc., it is possible to localize the disintegration points of the positrons.
It can be readily understood that this display is only obtained for one section of the organ 10 situated inside the plane of the ring 12 of the detector 14.
So as to obtain a volume display, several rings 12 are juxtaposed next to one another.
There are known means to improve the sensitivity of positron tomographs by carrying out a complementary measurement, known as a flight measurement, of the gamma radiations which makes it possible to localize each disintegration point of a positon on the corresponding track. The time of flight corresponds to the arrival time difference of two gamma radiations coincidence-detected. A knowledge of the position of the activated detectors 14 obtained by detection of the tracks, the speed of the light and the time of flight makes it possible to localize the emission site of the gamma radiations. In fact, the temporal resolution of the detectors 14 and of the associated analysis electronics only allows for a partial resolution of about 5 cm in the current state of the prior art. The localization obtained by this means is thus not sufficiently precise so as to draw up a cartograph of the disintegration points; it simply makes it possible to restrict the reconstruction range to the track portion situated inside the zone delimited by the time of flight and consequently does not take into account the noise appearing outside this zone. Owing to its inadequate temporal resolution, the time of flight tomograph does not make it possible to dispense with image reconstruction techniques.
Furthermore, in a conventional positron tomograph constituted by one or several rings of discrete detectors, the spatial resolution at the center of the device is roughly equal to the half-width of a detector. Now, a track is separated from its neighboring track by the step between two detectors, which is clearly at least equal to one detector width. The result is that the information received is subsampled. So as to overcome this drawback, the ring or rings is/are driven by an alternating rotary movement with low amplitude enabling each detector to occupy all the intermediate positions between its position at rest and those of its neighboring detectors. This small cicular movement with a radius equal to one half-step between detectors is commonly referred to as "wobbling".
This wobbling required for the sound functioning of positron tomographs results in a significant mechanical complexity rending it difficult to implement the device and proves to be expensive.
There is also another device shown in FIG. 2, called the "gamma camera" known as SPECT (Single Photon Emission Computed Tomograph) allowing for the display of gamma radiations emitted by an organ rendered radio-emitting by injecting a gamma tracer which only produces one gamma radiation per disintegration. There is a description of this SPECT in the patent U.S. Pat. No. 3,011,057. Contrary to the case with positron tomographs which embody an electronic collimation, the SPECTS use a material collimator 16. This collimator is formed by a plate made of an absorbant material, such as lead, pierced with holes orientated either parallel to each other and perpendicular to the surface of the scintillator 18 to which the collimator 16 is attached, or in a direction converging towards the center of the device so as to obtain an enlarging effect. The size and orientation of the holes are such that only the gamma radiations having a corresponding propagation direction are transmitted.
The gamma radiations interacting with the scintillator 18 provoke scintillations which are detected by several photoelectron multiplier tubes 22 disposed side by side on the rear face of the scintillator. The photoelectron multiplier tubes are distanced from the rear face of the scintillator 18 by a film 20 made of a transparent material.
If the scintillator 18 used was in direct contact with the photoelectron multipliers, a scintillation occuring at the center of the detection zone of a photoelectron multiplier and extremely close to the latter could only be detected with difficulty by other photoelectron multipliers without this transparent film 20. Now, it is important that each scintillation is detected by several photoelectron multipliers 22.
In fact, one barycentric measurment of the position of each scintillation is effected by electronic means 24 with the aid of signals proportional to the light intensity detected, these signals being delivered by the photoelectron multiplier tubes 22.
These electronic barycentric measurement means 24 connected to the outputs of the photoelectron multipliers 22 deliver signals indicating the average position of each scintillation on the scintillator 18. This position is solely defined along two directions contained in a plane parallel to the coupling face between the scintillator and the photoelectron multiplier.
So as to obtain a three-dimensional image of the organ, the SPECT rotates around the organ: thus, the directions of the gamma radiations are detected.
All the directions are processed by an image and display reconstruction system (not shown) making it possible to obtain an image of the points emitting a gamma radiation in the organ studied.
Owing to the need to obtain gamma radiations with specific directions before they interact with the scintillator 18, the collimator 16 is required to have holes with an extremely reduced diameter. The quantity of gamma radiations giving rise to scintillations is extremely small. In other words, the sensitivity of a SPECT is slight, which constitutes the major drawback of this type of device. In fact, the lower the sensitivity of the detector is, the stronger needs to be the dose of the radioactive element injected into the organ, this dose being limited when a living organ is to be marked.
In positron tomographs as well as in SPECTS, so as to obtain a three-dimensional image with acceptable resolution, it is necessary to carry out an image reconstruction treatment, which considerably adds to the complexity of the devices.